1. Field of the Invention
The present invention pertains generally to the use of a detachable x-ray detector array in combination with an existing radiation therapy x-ray simulator to produce computed tomographic (CT) image reconstructions. In more detail, the present invention relates to an apparatus for producing a CT scan from the width-collimated fan beam produced by an existing x-ray simulator and to a method for transforming the data produced by the detector array of the simulator into a back projected image of the target object. More particularly, this invention pertains to a method for reconstructing x-ray projection data to form CT images from divergent x-ray beams. The method can utilize noisy x-ray projection data, for example, data resulting from poor x-ray power supplies, and correct for geometrical errors resulting from mechanical defects in the motion of the beam projector that cause the central projection ray to move about the center of the detector array as a function of the angle of beam projection.
2. Introduction
X-ray computed tomography (CT) is a technique for obtaining cross-sectional reconstructions of three dimensional objects using x-rays. In the simplest example of CT imaging, a narrow beam of penetrating x-rays is scanned across an object or patient in synchrony with a radiation detector on the opposite side of the patient. If the beam is monoenergetic or nearly so, the transmission of x-rays through the patient is given by the equation EQU I=I.sub.0 exp(-.mu.x) [1]
where the patient is assumed to be a homogeneous medium with the attenuation coefficient .mu.. If the x-ray beam is intercepted by two regions with attenuation coefficients .mu..sub.1 and .mu..sub.2 and thicknesses x.sub.1 and x.sub.2, the x-ray transmission is characterized as EQU I-I.sub.0 exp-(.mu..sub.1 x.sub.1 +.mu..sub.2 x.sub.2) [2]
This formula is generalized to many (n) regions with different linear attenuation coefficients with the argument of the exponent ##EQU1## Separate attenuation coefficients cannot be determined with a single transmission measurement because there are too many unknown values of .mu..sub.i in the equation. However, with multiple transmission measurements at different orientations of the x-ray source and detector, the separate coefficients can be distinguished so that a cross-sectional display of coefficients is obtained across the plane of transmission measurements. By assigning gray levels (see below) to different ranges of attenuation coefficients, a display is obtained that represents various structures in the patient with different x-ray attenuation characteristics. This gray scale display of attenuation coefficients constitutes a CT image.
The numbers computed by the reconstruction algorithm are not exact values of attenuation coefficients. Instead, they are integrals, termed CT numbers, which are related to attenuation coefficients. On most newer CT units, the CT numbers range from -1,000 for air to +1,000 for bone, with the CT number for water set at 0. CT numbers normalized in this manner are termed Hounsfield units and provide a range of several CT numbers for a one percent (1%) change in attenuation coefficients.
To portray the CT numbers as a gray scale visual display, a CRT is used. The CRT includes a contrast enhancement feature that superimposes the shades of gray available in the display device (i.e., the dynamic range of the display) over the range of CT numbers of diagnostic interest. Control of image contrast with the contrast enhancement feature is essential in x-ray computed tomography because the electron density, and therefore the x-ray attenuation, is remarkably similar for most tissues of diagnostic interest. These electron densities vary from 3.07.times.10.sup.23 elec/cc for fat tissue to 5.59.times.10.sup.23 elec/cc for the densest tissue, bone. Lung tissue has a much lower electron density, 0.83.times.10.sup.23 elec/cc, because of the alveolar and bronchial spaces.
The first CT systems were introduced in approximately 1971 by the EMI Corporation of England. These early systems used an x-ray source mounted in a gantry with detectors. The patient was inserted between the x-ray source and the detectors and the joined x-ray source and detectors were moved about the patient to obtain projection rays through the patient. These values were fed to a computer which then reconstructed a cross sectional image of the plane through which the pencil beam of x-rays passed. During this translational scan of perhaps 40 cm in length, multiple (e.g., 160) measurements of the x-ray transmission were obtained. Next, the angular orientation of the scanning device was incremented one degree and a second translational scan of 160 transmission measurements was performed. This process was repeated at one degree increments through an arc of 180 degrees so that 28,800 x-ray transmission measurements were accumulated. Those measurements were then transmitted to a computer equipped with a mathematical algorithm for reconstructing an image of attenuation coefficients across the anatomical plane defined by the scanning x-ray beam.
Although this approach yielded satisfactory images of stationary objects, considerable time (4-5 minutes) was required for data accumulation and the images were subject to motion blurring. Soon after the introduction of pencil beam scanners, fan-shaped x-ray beams were introduced so that multiple measurements of x-ray transmission could be made simultaneously. Fan beam geometries, with increments of a few degrees for the different orientations (e.g., a 30-degree fan beam and 10-degree angular increments), reduced the scan time to 20-60 seconds and improved the image quality by reducing the effects of motion. Computed tomographic scanners with x-ray fan beam geometries and multiple radiation detectors constituted the second generation of CT scanners.
In late 1975, the third generation of CT scanner was introduced. These scanners eliminated the translational motion of previous scanners, using rotational motion of the x-ray tube and detector array or rotational motion of the x-ray tube within a stationary circular array of 600 or more detectors. With these scanners, data accumulation times as fast as two seconds are achievable.
Both stationary and rotating anode x-ray tubes are used in CT scanners. Many of the translation-rotation CT scanners have an oil-cooled, stationary anode x-ray tube with a focal spot on the order of 2.times.16 mm. The limited output of these x-ray tubes necessitates a sampling time of about 5 msec for each measurement of x-ray transmission. This sampling time, together with the time required to move and rotate the source and detector, limits the speed with which data can be accumulated with CT units using translational and rotational motion.
To reduce the sampling time of 2-3 msec, most fast-scan CT units use rotating-anode x-ray tubes, often with a pulsed x-ray beam, to achieve higher x-ray outputs. Even with rotating-anode tubes, the heat-storage capacity of the anode may be exceeded if cooling periods are not observed between sets of successive images.
After transmission through the patient, the x-ray beam is collimated to confine the transmission to a slice with a thickness of a few millimeters and to reduce scattered radiation to less than one percent (1%) of the primary beam intensity. The height of the collimator defines the thickness of the CT slice. This height, when combined with the area of a single picture element (pixel) in the display, defines the three-dimensional volume element (voxel) in the patient corresponding to the two-dimensional pixel of the display. A voxel encompassing a boundary between two tissue structures (e.g., muscle and bone) yields an attenuation coefficient of the two structures. This "partial volume artifact" may be reduced by narrowing the collimator to yield thinner slices. However, this approach reduces the intensity of the x-rays incident upon the detector and the detector signals are subject to greater statistical fluctuations, thus introducing more noise into the displayed image.
To reduce the detector response time, all detectors used in CT scanning are operated in current rather than pulse mode. Also, rejection of scattered radiation is assigned to the detector collimator rather than to pulse height analyzers. Detectors for CT scanning are chosen on the basis of detection efficiency (greater than 50%), short response time and stability of operation, and are either gas-filled ionization chambers or solid scintillation detectors. With any detector, the stability of response from one transmission measurement to the next is essential for the production of artifact-free reconstruction images. With a pure rotational source and detector geometry, for example, detector instability gives rise to ring-shaped artifacts in the image. Minimum energy dependence of the detector over the energy range for the CT x-ray beam also is important if corrections for beam hardening are to be applicable to all patient sizes and configurations.
All of the early CT systems were designed and built only to perform CT studies. The concept of using other types of radiation sources that had not been specifically designed for CT imaging was initiated in the mid 1970's.
Several of these efforts utilized existing x-ray therapy simulators. An x-ray simulator is a device that duplicates a radiation treatment unit in terms of its geometric, mechanical and optical properties, but uses a diagnostic x-ray tube as the source of radiation to simulate the properties of the treatment beam. A simulator allows the beam direction and the treatment fields to be determined while encompassing the target object with the simulator's irradiation. Since the simulator's emissions are generally less intense and less energetic than the emissions of therapy devices, there is a reduction in the target object's exposure to radiation.
The combination of a detector system and an x-ray therapy simulator provides the necessary front end of a CT system. Application of the requisite information processing techniques and algorithmic reconstruction processes in combination with the simulator cum detector system enable production of CT images.
In another type of radiography, known as panoramic radiography, the x-ray source and detector lie in the same plane just as is the case for an x-ray simulator, but they are rotated in a plane that is transverse to the plane containing the source and detector. In panoramic x-ray devices such as those used to produce images of the maxillo-facial region, the center of rotation of the source and detector moves along an elliptical path, the output signal of the detector being read periodically and then conditioned to produce a panoramic x-ray image of the target object. In contrast with a CT scanner, such devices typically scan only approximately 230.degree.-240.degree. about the target object.